The present invention pertains generally to ophthalmic diagnostic equipment. More particularly, the present invention pertains to devices and methods for examining the fundus of the eye. The present invention is particularly, but not exclusively useful for generating an aberration-free beam of light that can be focused into the fundus with a focal depth of approximately twenty microns.
Using an optical device to examine human tissue requires the ability to precisely focus an incident light beam to a predetermined focal spot on the tissue. This can be difficult to accomplish. In the specific case of the fundus of an eye (i.e. the retina), the size of the focal spot that is required for an effective examination is infinitesimal. The task is further complicated by the fact that the light beam must be effectively directed through the eye before it is incident at its focal spot in the fundus.
The anatomy of the fundus of an eye is known to comprise several distinct layers. These layers are of different types of tissue, and in an anterior to posterior direction they include: axons, ganglion cells, bipolar cells, receptors (rods and cones), pigment cells and the choroid. Together, these layers establish a depth for the fundus that is somewhere in the range of about three hundred and fifty microns, plus or minus a hundred microns (350 xcexcmxc2x1100 xcexcm). Individually, however, some layers are only about twenty microns (20 xcexcm) in depth (e.g. ganglion cells and pigment cells). Therefore, in order to effectively examine the various layers of the fundus, it is necessary to resolve a volume of tissue having a depth that is equal to or less than 20 xcexcm. Stated differently, this requires an ability to focus incident light onto a focal pot in the fundus that has a three dimensional point spread function (PSF), i.e. the finest volume to be resolved, that has a twenty micron depth. Optical physics and the anatomical structure of the eye, however, affect the ability to do this.
Whenever a beam of light is focused to a point, it happens that the light beam is cone shaped in the space between the focusing element (e.g. a lens) and the point to which the beam is focused (i.e. a focal spot). In accordance with optical physics, it also happens that the depth of focus at the focal spot will decrease dramatically as the cone angle of the focused light is increased. As indicated above, for fundus examinations, it is desirable to have as shallow a depth for the focal spot as is possible. Consequently, there is a need for relatively large cone angles. For the human eye, however, the cone angle for light that can be focused onto the fundus is limited in at least two ways. For one, the iris of the eye limits the amount of light that can enter the eye. At most, the iris can be dilated to establish an aperture that is only about a six millimeters (6 mm) in diameter. For another, the cornea and lens of the eye introduce significant optical aberrations into the light that is being focused onto the fundus when the aperture is increased beyond about two millimeters (2 mm).
To contrast the consequences for an optical examination of the fundus of a normal eye, it is helpful to appreciate the difference between the possible PSFs for a 2 mm aperture and a 6 mm aperture. With a 2 mm aperture, the PSF is a volume that is approximately 10 xcexcmxc3x9710 xcexcmxc3x97200 xcexcm. On the other hand, with a 6 mm aperture, a PSF with a volume of 2 xcexcmxc3x972 xcexcmxc3x9720 xcexcm, is possible. The assumption here has been that the light being focused is aberration-free. Thus, for optical examinations of the fundus, wherein the depth of the focal spot needs to be limited to about 20 xcexcm, it is very desirable to be able to use the full 6 mm potential for the aperture of the iris. Additionally, due to differences in the tissues of the retina that are to be imaged, it is desirable to use different imaging modalities. Specifically, blood in tissues of the retina need to be imaged using techniques that are different from the techniques used to image essentially transparent tissue.
In light of the above, it is an object of the present invention to provide a device for imaging the fundus of the human eye which is capable of generating an essentially aberration-free beam of light for use in diagnosing the fundus. It is another object of the present invention to provide a device for imaging the fundus that can effectively focus light into the fundus with a focal depth (i.e. PSF) of about twenty microns. Still another object of the present invention is to provide a device for imaging the fundus of the human eye which is capable of examining individual layers of tissue in the fundus using appropriate imaging modalities. Yet another object of the present invention is to provide a device for imaging the fundus of the human eye that is easy to use, relatively simple to manufacture and comparatively cost effective.
In accordance with the present invention, a device for aberration-free imaging of the fundus of the human eye includes, in part, a plurality of imaging units and an active mirror. The imaging units are respectively used for viewing the fundus (i.e. retina) of the eye with different imaging modalities, and the active mirror can be programmed to effectively remove aberrations from the light that is focused onto the fundus and which is subsequently reflected from the fundus and received by the imaging unit. In more detail, the device of the present invention includes a light source for generating a light beam, and there is an optical element for focusing the light beam through the eye to a focal spot in the fundus. The light that is reflected from the focal spot can then be imaged. Depending on the imaging modality, different imaging units will be used to do this. For example, the imaging unit may be a fluoroscope if blood vessels in the retina are to be imaged. On the other hand, the imaging unit may be an ellipsometer if transparent tissue is to be imaged.
Regardless the imaging modality being used, when the light beam is reflected from the focal spot in the fundus, it will exhibit a reflected wavefront that is characteristic of the reflected beam. This reflected wavefront is then directed, by the active mirror, toward the imaging unit. Due to aberrations introduced by the eye, however, the reflected wavefront requires some compensation before it reaches the imaging unit.
In order to compensate the wavefront that is reflected from the fundus of the eye, a computer/comparator is used to compare the reflected wavefront with an aberration-free wavefront (e.g. a plane wavefront). Specifically, this comparison is made to establish an error signal. The computer/comparator then uses the error signal to program the active mirror to accomplish two separate, but interrelated functions. For one, the programmed active mirror compensates for aberrations that are introduced into the light beam before it is incident on the focal point in the fundus. For another, the programmed active mirror compensates the reflected wavefront, to thereby create an aberration-free wavefront that is received by the imaging unit.
It is an important aspect of the present invention that the eye be as widely dilated as possible when the light beam is focused to a focal spot in the fundus and, accordingly, when the light is reflected from the focal spot and received at the imaging unit. Specifically, the aperture of the eye during the operation of the present invention is preferably dilated to about six millimeters in diameter. Normally, the aperture is around two millimeters in diameter. The consequence of this dilation is two-fold. First, the extended aperture (i.e. 6 mm) allows for a much improved point spread function (PSF) for the focal spot. Specifically, PSF is a three dimensional measurement that defines the finest volume of focus for a light beam. It happens that when the aperture of the eye is dilated to about 6 mm, the PSF can be as small as about 2 xcexcmxc3x972 xcexcmxc3x9720 xcexcm. In contrast, with a two millimeter diameter aperture for the eye, the PSF is more on the order of 10 xcexcmxc3x9710 xcexcmxc3x97200 xcexcm. Second, with an extended aperture, significant aberrations are introduced into light as it passes into and out of the eye through the cornea. Nevertheless, in order to effectively image individual layers with the fundus of the eye that may be as shallow as 20 xcexcm, it is necessary to have the improved PSF with a depth of focus that is around 20 xcexcm. Thus, the purpose of the present invention is to compensate for the introduced aberrations and, consequently, be able to benefit from the improved PSF.
As envisioned for the present invention, the imaging unit can be any device or apparatus known in the art, such as an ellipsometer or a fluoroscope. Also, the light source can be of any type known in the art, such as a laser diode or a super luminescence diode (SLD).